The phenomenon of soliton self-frequency shift (SSFS) in optical fiber in which Raman self-pumping continuously transfers energy from higher to lower frequencies (Dianov et al., JETP. Lett. 41:294 (1985)) has been exploited over the last decade in order to fabricate widely frequency-tunable, femtosecond pulse sources with fiber delivery (Nishizawa et al., IEEE Photon. Technol. Lett. 11:325 (1999); Fermann et al., Opt. Lett. 24:1428 (1999); Liu et al., Opt. Lett. 26:358 (2001); Washburn et al., Electron. Lett. 37:1510 (2001); Lim et al., Electron. Lett. 40:1523 (2004); Luan et al., Opt. Express 12:835 (2004). Because anomalous (positive) dispersion (β2<0 or D>0) is required for the generation and maintenance of solitons, early sources which made use of SSFS for wavelength tuning were restricted to wavelength regimes greater than 1300 nm where conventional silica fibers exhibited positive dispersion (Nishizawa et al., IEEE Photon. Technol. Lett. 11:325 (1999); Fermann et al., Opt. Lett. 24:1428 (1999)). The recent development of index-guided photonic crystal fibers (PCF) and air-core photonic band-gap fibers (PBGF) relaxed this requirement with the ability to design large positive waveguide dispersion and therefore large positive net dispersion in optical fibers at nearly any desired wavelength (Knight et al., IEEE Photon. Technol. Lett. 12:807 (2000)). This allowed for a number of demonstrations of tunable SSFS sources supporting input wavelengths as low as 800 nm in the anomalous dispersion regime (Liu et al., Opt. Lett. 26:358 (2001); Washburn et al., Electron. Lett. 37:1510 (2001); Lim et al., Electron. Lett. 40:1523 (2004); Luan et al., Opt. Express 12:835 (2004)).
Unfortunately, the pulse energy required to support stable Raman-shifted solitons below 1300 nm in index-guided PCFs and air-core PBGFs is either on the very low side, a fraction of 1 nJ for silica-core PCFs, (Washburn et al., Electron. Lett. 37:1510 (2001); Lim et al., Electron. Lett. 40:1523 (2004)) or on the very high side, greater than 100 nJ (requiring an input from an amplified optical system) for air-core PBGFs (Luan et al., Opt. Express 12:835 (2004)). The low-energy limit is due to high nonlinearity in the PCF. In order to generate large positive waveguide dispersion to overcome the negative dispersion of the material, the effective area of the fiber core must be reduced. For positive total dispersion at wavelengths less than 1300 nm this corresponds to an effective area, Aeff, of 2-5 μm2, approximately an order of magnitude less than conventional single mode fiber (SMF). The high-energy limit is due to low nonlinearity in the air-core PBGF where the nonlinear index, n2, of air is roughly 1000 times less than that of silica. These extreme ends of nonlinearity dictate the required pulse energy (U) for soliton propagation, which scales as UD·Aeff/n2. In fact, most microstructure fibers and tapered fibers with positive dispersion are intentionally designed to demonstrate nonlinear optical effects at the lowest possible pulse energy, while air-core PBGFs are often used for applications that require linear propagation, such as pulse delivery. For these reasons, previous work using SSFS below 1300 nm were performed at soliton energies either too low or too high (by at least an order of magnitude) for many practical applications, such as multiphoton imaging where bulk solid state lasers are currently the mainstay for the excitation source (Diaspro, A., Confocal and Two-Photon Microscopy, Wiley-Liss: New York (2002)).
Applications of Femtosecond Sources in Biomedical Research.
There are a number of biomedical applications that require femtosecond sources. Although applications requiring a large spectral bandwidth (such as optical coherence tomography) can also be performed using incoherent sources such as superluminescent diodes, techniques based on nonlinear optical effects, such as multiphoton microscopy and endoscopy, almost universally require the high peak power generated by a femtosecond source.
Endoscopes play an important role in medical diagnostics by making it possible to visualize tissue at remote internal sites in a minimally invasive fashion. The most common form employs an imaging fiber bundle to provide high quality white light reflection imaging. Laser scanning confocal reflection and fluorescence endoscopes also exist and can provide 3D cellular resolution in tissues. Confocal endoscopes are now becoming available commercially (Optiscan Ltd, Australia, Lucid Inc, Rochester) and are being applied in a number of clinical trials for cancer diagnosis. Multiphoton excitation based endoscopes has attracted significant attention recently. There were a number of advances, including fiber delivery of excitation pulses, miniature scanners, double clad fibers for efficient signal collections, etc. Thus, just like multiphoton microscopes have proven to be a powerful tool in biological imaging, multiphoton endoscopes have great potentials to improve the capability of the existing laser-scanning optical endoscopes. It is quite obvious that a compact, fully electronically controlled, femtosecond system seamlessly integrated with fiber optic delivery is essential for multiphoton endoscopy in medical diagnostics, particularly to biomedical experts who are not trained in lasers and optics.
Perhaps the most promising and successful area in biomedical imaging that showcases the unique advantage of multiphoton excitation is imaging deep into scattering tissues. In the past 5 to 10 years, multiphoton microscopes have greatly improved the penetration depth of optical imaging and proven to be well suited for a variety of imaging applications deep within intact or semi-intact tissues, such as demonstrated in the studies of neuronal activity and anatomy, developing embryos, and tissue morphology and pathology. When compared to one-photon confocal microscopy, a factor of 2 to 3 improvement in penetration depth is obtained by multiphoton microscopes. Nonetheless, despite the heroic effort of employing energetic pulses (˜μJ/pulse) produced by a regenerative amplifier, multiphoton microscopes have so far been restricted to less than 1 mm in penetration depth. One promising direction for imaging deep into scattering tissue is to use longer excitation wavelength. Although the “diagnostic and therapeutic window,” which is in between the absorption regions of the intrinsic molecules and water, extends all the way to approximately 1300 nm, previous investigations involving multiphoton imaging are almost exclusively carried out within the near IR spectral window of approximately 0.7 to 1.1 μm, constrained mostly by the availability of the excitation source. Currently, there are only two femtosecond sources at the spectral window of 1200 to 1300 nm, the Cr:Forsterite laser and the optical parametric oscillator (OPO) pumped by a femtosecond Ti:Sapphire (Ti:S) laser. In terms of robustness and easy operation, both sources rank significantly below the Ti:S laser. Thus, the development of a reliable fiber source tunable from 1030 to 1280 nm will open up new opportunities for biomedical imaging, particularly for applications requiring deep tissue penetration.
Femtosecond Sources for Multiphoton Imaging.
Shortly after the inception of multiphoton microscopes, mode-locked solid state femtosecond lasers, most commonly the Ti:S lasers, have emerged as the favorite excitation sources to dominate the multiphoton microscope field today. When compared to earlier ultrafast lasers, e.g., ultrafast dye lasers, the Ti:S lasers are highly robust and flexible. The concurrent development of the mode-locked Ti:S lasers was perhaps the biggest gift for multiphoton microscopes and enabled them to rapidly become a valuable instrument for biological research. Nonetheless, the cost, complexity, and the limited potential for integration of the bulk solid state lasers have hampered the widespread applications of multiphoton microscopes in biological research. The fact that a disproportionate number of multiphoton microscope systems are located in physics and engineering departments, instead of the more biologically oriented institutions, reflects at least in part the practical limitations of the femtosecond pulsed source. Obviously, the requirement of a robust, fiber delivered, and cheap source is even more urgent for multiphoton endoscopy in a clinical environment.
Mode-locked femtosecond fiber lasers at 1.03 and 1.55 μm have been improving significantly in the last several years, mainly in the output pulse energy (from 1 to ˜10 nJ). Even higher pulse energy can be achieved in femtosecond fiber sources based on fiber chirped pulse amplification. However, femtosecond fiber sources, including lasers and CPA systems, have seen only limited applications in multiphoton imaging. The main reason is that they offer very limited wavelength tunability (tens of nanometers at best), severely restricting the applicability of these lasers, making them only suitable for some special purposes. In addition, existing femtosecond fiber sources at high pulse energy (>1 nJ) are not truly “all fiber,” i.e., the output are not delivered through a single mode optical fiber. Thus, additional setup, typically involving free-space optics, must be used to deliver the pulses to the imaging apparatus, partially negating the advantages of the fiber source. Reports have demonstrated the possibility of propagating femtosecond IR pulses through a large core optical fiber at intensities high enough (˜1 nJ) for multiphoton imaging. In addition, a special HOM fiber that is capable of delivery energetic femtosecond pulses (˜1 nJ) has been demonstrated. However, both fibers have normal dispersion, and both require a free-space grating pair for dispersion compensation. Not only is such a grating pair lossy and complicated to align, it needs careful adjustment for varying fiber length, output wavelength, and output pulse energy, and falls short of the requirement for most biomedical research labs and future clinical applications.
The present invention is directed to overcoming these and other deficiencies in the art.